1. Field
The embodiments described herein relate generally to radiation-based imaging. More particularly, the described embodiments relate to the estimation of scatter radiation within projection images.
2. Description
Radiation-based imaging systems are well-known. According to some examples, a radiation beam is emitted prior to treatment, passes through a volume of the patient and is received by an imaging system. The imaging system produces a set of data that represents the attenuative properties of objects of the patient volume that lie between the radiation source and the imaging system.
The set of data is used to generate a two-dimensional projection image of the patient volume. The projection image will include areas of different intensities that reflect different compositions of the objects. For example, areas of low radiation intensity may represent bone and areas of high radiation intensity may represent tissue. Several two-dimensional projection images may be acquired from different perspectives with respect to the patient volume and combined to generate a three-dimensional image of the patient volume.
The three-dimensional image, and “slices” thereof, may be used to diagnose disease, to guide surgical interventions, to plan radiation therapy, to confirm patient positioning prior to therapy, and/or to perform image-guided radiotherapy (IGRT). Advanced clinical applications such as dose-guided radiation therapy (DGRT) rely on images in which displayed Hounsfield numbers are accurately mapped to electron density to ensure the accuracy of calculated doses.
Conventional computed tomography (CT) reconstruction implicitly assumes that the imaging beam is monoenergetic, and, consequently, that the acquired projection data are related to the attenuation coefficients within the imaged object by a simple exponential model. Owing to the polyenergetic nature of Bremsstrahlung X-ray beams, the scattering of photons within the imaged object, and the dependence of detector response on incident photon energy, this assumption often leads to large inaccuracies. In addition, the intensity and energy spectrum of cone beam radiation sources typically varies over the imaging field. Accurate CT reconstruction may therefore depend on appropriate modeling of scatter, beam-hardening, beam profile, beam spectral variation and detector response.
Scatter radiation, for example, does not generally travel along an expected radiation trajectory from the radiation source to the imaging system. Scatter radiation received at a particular location of the imaging system therefore does not reflect attenuative properties of all the tissues located along an expected trajectory from the radiation source to the particular location. This leads to quantitative inaccuracies in reconstructed tomographic images. In medical X-ray imaging, these inaccuracies manifest themselves as dark shading or streak artifacts, where the density of the images object is underestimated owing to scatter and beam-hardening. In addition, received scatter radiation induces noise and reduces the intensity gradients (i.e., contrast) between image areas that represent different objects in a projection image. The reduced contrast may inhibit identification of structures within the projection image and any CT reconstruction based thereon, particularly with respect to soft tissue structures.
Many methods of scatter reduction and compensation have been proposed. Most methods are based on simplified models of scatter physics or assumptions relating to the smoothness of the scatter distribution in the projection images. Full Monte Carlo (MC) simulation methods are too time consuming, at present, for practical use in many clinical workflows and require an existing dataset describing the attenuation coefficient distribution within the object. To address these limitations, iterative schemes have been proposed whereby MC simulations are applied to a tomographic image reconstructed from scatter-contaminated projections, and subsequently to scatter-corrected iterates thereof. While such schemes require multiple MC runs, the application of variance reduction methods such as smoothing allows a low number of particle histories to be employed.
Beam-stop arrays can be used to measure scatter directly, but require the acquisition of two sets of projection images. Other beam-stop methods interpolate scatter between collimator shadows. These latter methods result in a reduction in field-of-view and performance that is strongly object-dependent. Antiscatter grids can physically attenuate scatter radiation by 5 to 10 times relative to primary radiation but lead to the loss of 25% to 33% of primary radiation and reduced flexibility with regard to detector positioning. Focused antiscatter grids (Potter-Bucky grids) can be employed for megavoltage (MV) X-ray and 60Co gamma ray imaging, but are bulky, heavy and costly. Scatter radiation may also be separated by physical frequency modulation and filtering of the scatter and primary radiation components, but this is difficult to achieve with polyenergetic beams.
The scatter kernel superposition (SKS) method of scatter modeling has been developed for application to MV imaging using a treatment linear accelerator (linac). To remove the contribution of scattered X-rays from a projection image, SKS methods assume: 1) the scatter distribution may be modeled as the sum of the scatter contributions of primary pencil beams that traverse the imaged object; 2) the scatter contribution of each primary pencil beam is dependent only on the material it traverses along its path; and 3) a primary image may be recovered from the measured (primary plus scatter) image via an iterative process of scatter removal.
The first two assumptions are most valid in homogeneous objects but are violated in the case of heterogeneous objects and/or long, thin objects. In heterogeneous objects, scattered particles may follow radiological paths through materials having properties that differ markedly from those traversed by the primary pencil beam. When long, thin objects are imaged (i.e., with the long axis parallel to the beam), SKS methods will calculate an equal contribution for two parallel pencil beams, one traversing the central axis of a cylinder and the other traversing the edge, even though scatter due to the beam at the edge experiences strongly anisotropic attenuation.
The effects of violating these assumptions are increased at low energies. A given anthropomorphic phantom will appear far less homogeneous to an incident beam at diagnostic (i.e., kV) energies owing to the far greater dependence of attenuation on the atomic number of the imaged material. Additionally, the increased attenuation experienced by a kV primary beam leads to higher scatter-to-primary ratios (SPRs). The higher SPRs increase errors due to the above-described problems with long thin objects, since differential attenuation of the scatter distribution attributable to the two central and edge-located pencil beams is larger. Moreover, as will be further described below, the wider scattering angles of low energy scattering events lead to wider scatter kernels and thus more extended propagation of estimation errors at object edges.